The present invention relates to the efficient utilization of power in body tissue stimulators, and more particularly to an improved power control loop for implantable cochlear stimulator systems. Such implantable cochlear stimulator systems provide improved hearing for the hearing impaired. The power control loop serves the important function of providing power to the implanted part of the cochlear stimulator system and the efficiency of the power control loop is critical in developing the miniaturized systems of the future.
U.S. Pat. No. 4,400,590 issued Aug. 23, 1983 for xe2x80x9cApparatus for Multi-Channel Cochlear Implant Hearing Aid Systemxe2x80x9d describes and illustrates a system for electrically stimulating predetermined locations of the auditory nerve within the cochlea of the ear, which system includes a multi-channel intra-cochlear electrode array. The electrode array comprises a plurality of exposed electrode pairs spaced along and embedded in a resilient curved base for implantation in accordance with the method of surgical implantation described in U.S. Pat. No. 3,751,605 issued Aug. 7, 1973 for xe2x80x9cMethod of Inducing Hearing.xe2x80x9d The hearing aid system described in the ""590 patent receives audio signals at a signal processor located outside the body of a hearing impaired patient. The processor converts the audio signals into analog data signals which are transmitted by a cable connection through the patient""s skin to the implantable multi-channel intra-cochlear electrode array. The analog signals are applied to selected ones of the plurality of exposed electrode pairs included in the intra-cochlear electrode array to electrically stimulate predetermined locations of the auditory nerve within the cochlea of the ear where the intra-cochlear electrode array is positioned.
The cochlea stimulating system described in the ""590 patent is limited in the number of channels of information, the speed of transfer of stimulating signals to the cochlea and the fidelity of the signals. Also, the cable connection through the skin of the patient to the intra-cochlear electrode array is undesired in that it interferes with the freedom of movement of the patient and represents a possible source of infection.
U.S. Pat. No. 4,532,930, issued Aug. 6, 1985 for xe2x80x9cCochlear Implant System For an Auditory Prosthesisxe2x80x9d describes and illustrates a multiple electrode system which does not employ a through the skin connector. While multiple electrodes are employed to stimulate hearing, the system only operates with a single pulsatile output stimulating a single electrode channel at any given time. Such a sequential system is limited in speed of operation, and does not provide for analog operation where continuous stimulating signals, controllable in amplitude, are simultaneously applied to a number of electrode channels. Further, the system is subject to charge imbalance with misprogramming or circuit fault and inefficient use of electrical power. Moreover, once the stimulator unit is implanted there are no means for monitoring its ongoing circuit operation or power requirements so as to optimize its continued operation.
U.S. Pat. No. 4,592,359, issued Jun. 3, 1986 for xe2x80x9cMulti-Channel Implantable Neural Stimulatorxe2x80x9d describes a cochlear implant system having 4 current sources and 4 current sinks per channel, controlled by series switches, to provide 16 different circuits for supplying 16 levels of 2 polarities to each output channel. In a pulsatile mode, the system provides for simultaneous update (amplitude control) and output to all channels. However, the system does not permit simultaneous analog update and output on all channels and the electrode pairs for each channel are not electrically isolated from all other electrode pairs whereby undesired current leakage may occur. Further, once the stimulator is implanted there are no means for monitoring its ongoing circuit operation or power requirements so as to optimize its continued operation.
U.S. Pat. No. 4,947,844, issued Aug. 14, 1990 for xe2x80x9cReceiver/Stimulator For Hearing Prosthesisxe2x80x9d describes and illustrates a multiple channel electrode system. The system includes an implantable receiver/stimulator connected to an implantable electrode array. Included in the implantable receiver/stimulator is a transmitter for telemetering one electrode voltage, measured during stimulation, to an external receiver for monitoring and analysis as an indicator of proper operation of the implantable stimulator. The transmitter comprises an oscillator operating at a frequency of about 1 MHz. The output of the oscillator is coupled to the implant""s receiving coil. The oscillator signal, when received after transmission, is demodulated to recover the selected voltage waveforms. Unfortunately, such a telemetry system is not only limited to the monitoring of one voltage, but the simultaneous transmission of the telemetry signal and reception of the input carrier signal results in undesired modulation and possible loss of input data.
For cochlear stimulator applications, it is generally desirable to employ a cochlear stimulator that is driven by a behind-the-ear speech processor, e.g., of the type described in U.S. Pat. No. 5,824,022 issued Oct. 20, 1998 for xe2x80x98Cochlear stimulation system employing behind-the-ear speech processor with remote control.xe2x80x99 Behind-the-ear speech processors offer several advantages, but because of their small size are limited in the size of the battery they may carry (which in turn limits the useful life of the battery.) The small battery size results in a requirement for very low power dissipation. Although low power digital electronics have enabled digital hearing aids, this technology is only part of the answer for implantable stimulators. This is because an implantable stimulator, e.g., a cochlear stimulator, requires additional variable current to stimulate the target tissue, e.g., the auditory nerve within the cochlea, and this power must be transferred across a transcutaneous link, that is, at best, only about 50% efficient. While digital hearing aids only need to drive a transducer that uses less than one milliwatt (mW) of power, an implantable tissue stimulator may require up to 50 mW of stimulus power, which means (assuming a 50% transcutaneous link transfer efficiency) the need to transmit up to 100 mW of power to the implant device. Since power is proportional to the square of the voltage, it would thus be desirable to have a way to precisely and actively control the voltage in the implant device to track the output power requirements of the device. For example, for a cochlear stimulator, where room sound and speech levels are variant, it would be desirable to track speech and system variations and make automatic adjustments in the input power that track these variations, thereby only transmitting power to the implant device that is needed for the current conditions, thereby increasing the life of the battery.
U.S. Pat. No. 5,876,425, issued Mar. 2, 1999 for xe2x80x9cPower Control Loop for Implantable Tissue Stimulatorxe2x80x9d describes a feedback power control loop utilizing back telemetry from the implantable device. The Implantable Cochlea Stimulator (ICS) utilizes a tank capacitor as an internal rechargable power source. The ICS monitors the voltage level of the tank capacitor and back transmits the tank capacitor voltage to the Wearable Signal Receiver and Processor (WP). Based on the back transmitted tank capacitor voltage, the WP computes the power level to be transmitted to the ICS to maintain the tank capacitor voltage within acceptable levels. While the approach taught in the ""425 patent provides advantages over previous approaches that transmit power based on the peak ICS power requirement, it also results in delays in the calculation of the power requirements. The delays in the response of the power control loop result in too much power being transmitted at times, and this power is dissipated versus being stored. The requirement to continuously monitor the tank capacitor voltage requires that the implantable device expend power to digitize the measurement, and additional power is required to back transmit the tank capacitor voltage measurements. The power dissipation resulting from excessive power transmission and the power required to continuously provide tank capacitor voltage measurements, result in a requirement for either a larger battery in the WP, or more frequent recharging. The desire to develop a cochlear stimulator that is driven by a behind-the-ear speech processor further motivates the development of more efficient power control.
Accordingly, there is a continuing need for greater efficiency in supplying power to the implantable stimulator unit of the system to optimize system operation and power efficiency. The present invention satisfies such needs.
The present invention addresses the above and other needs by providing an active feed forward Power Control Loop (PCL) for implantable devices. Many implantable devices, and Implantable Cochlea Stimulator (ICS) systems in particular, require power to be supplied transcutaneously from an external device to an implantable device. A power supply method utilizing the wireless transmission of power from the external device to the implantable device is preferred over methods requiring a cable connection through the skin of the patient to the implantable device, because through the skin cables interfere with the freedom of movement of the patient and represents a possible source of infection. The use of a wireless power transmission from the external device to the internal device results in a requirement to control the amount of power transmitted by the external device.
Power for short term use is typically stored within the implantable device in a power storage device, for example, a tank capacitor. The tank capacitor is charged by RF power transmitted transcutaneously from the external device. The power source within the external device is generally a rechargeable battery. The size of the battery and the frequency of recharging is determined by the efficiency of the power use within the system. Due to the nature of the power storage within the implantable device, any power transmitted to the implantable device, in excess of that required to maintain the charge of the tank capacitor, is dissipated. As a result, optimal power utilization within the system requires that the power transmitted to the implantable device be a function of the instantaneous tank capacitor voltage.
The present invention advantageously utilizes a feed forward PCL, with a model of tank capacitor voltage as a function of the input speech level and the transmitted power level. A model of tank capacitor voltage is developed based on the implantable device circuitry, and as a function of externally measurable or controllable parameters. However, due to variable physical characteristics of the patient and the electronics (e.g. the RF power transmission efficiency or the impedance between electrodes,) the tank voltage model contains several unknown coefficients. The coefficients of the model are determined by operating the ICS system at various operating points for all of the independent parameters that drive the tank capacitor voltage. The values of all of the parameters and the resulting tank capacitor voltage are recorded during the patient fitting process as a set for each operating point. The total number of operating points is preferably several times the number of independent parameters. The results of the data collection are used as inputs to a multiple regression analysis. The multiple regression analysis results in a set of coefficients for the model of the tank capacitor voltage. During subsequent system operation, the tank capacitor voltage is estimated based on the model, and power is provided to the implantable device to maintain the desired tank capacitor voltage. Should the model drift with time, it may be periodically updated by back telemetry of the actual tank capacitor voltage.
In accordance with one aspect of the present invention, there is provided a timely estimate of the power requirements of the implantable device which results in a faster response to changes in the power requirements. By anticipating changes in the power requirements, the feed forward PCL reduces voltage downswings and subsequent voltage overshoot and thus provides a more efficient power utilization. The efficient use of the external device""s power permits the battery in the external device to be made as small and light as possible which is advantageous for the development of improved behind-the-ear speech processors.
It is an additional feature of the present invention to eliminate the requirement for continuous measurement of the tank capacitor voltage within the implantable device, and continuous back telemetry of the tank capacitor voltage from the implantable device to the external device. The requirement to continuously monitor the tank capacitor voltage requires that the implantable device expend power to digitize the measurement, and additional power is required to back transmit the tank capacitor voltage measurements. By avoiding most or all of the requirement to provide tank capacitor voltage to the WP, the power required by these two steps is avoided.